JPS6262383B2 - - Google Patents

Description

[Detailed description of the invention]

The present invention relates to an improvement in an image processing method and apparatus in a radiographic system used for medical diagnosis, and more specifically, the present invention relates to an improvement in an image processing method and apparatus in a radiographic system used for medical diagnosis. The present invention relates to an improvement of an image processing method and apparatus in a radiographic system in which radiographic image information is recorded on the recording material, and then this radiographic image information is read out and reproduced, and is recorded as a final image on a recording material. Radiation that has passed through the object is absorbed by a phosphor to record radiation image information, which is then scanned with a laser beam or the like to excite the emitted light, which is read by a photodetector. 2. Description of the Related Art Radiographic systems are known that record radiographic images on a recording medium such as photographic film by modulating a beam. (U.S. Pat. No. 3,859,527) This phosphor-based radiographic system is significantly superior to conventional silver halide radiographic systems in that it can record images over a wide radiation exposure range. Highly useful,
It is particularly useful as an X-ray photography system for the human body. On the other hand, X-rays are harmful to the human body when exposed to large amounts of radiation, so it goes without saying that it is desirable to obtain as much information as possible with a single X-ray photograph, but current X-ray photographic films are not suitable for photography. It is required to have both observation and interpretation aptitude,
Since the X-ray film is designed to satisfy each of these requirements to a certain degree, it cannot be said that the X-ray exposure range is sufficiently wide in terms of suitability for imaging, and the image quality of current X-ray photographic film is not always suitable for diagnosis. There was a problem that it could not be said to be sufficient. Furthermore, although the radiographic system using a phosphor disclosed in the aforementioned U.S. Pat. . In order to solve the above-mentioned problems, the present inventors have proposed a radiographic image processing method and apparatus that can simultaneously satisfy both suitability for imaging and suitability for observation and interpretation in a radiographic system using a phosphor. (Patent Application No. 53-163571) If you use this method and device,
Radiographic images with improved diagnostic performance can be obtained economically and at high speed. This method and apparatus Sorg the original image signal read out from the phosphor and generate an unsharp mask signal corresponding to very low spatial frequencies at each scanning point.
Sus, the emphasis coefficient is β, and the reproduced image signal is S′, the frequency component above the ultra-low spatial frequency is emphasized by performing the calculation expressed as S′=Sorg+β(Sorg−Sus). That is. In this method and device, the emphasis coefficient β
can be fixed or variable, and in the case of variable
It may be changed depending on the size of either Sorg or Sus. Here, the ultra-low spatial frequency means a spatial frequency of approximately 0.5 cycles/mm or less. However, according to the inventor's subsequent research, it was found that when β is fixed, artifacts are likely to occur in low and high brightness regions. On the other hand,
When β is made variable, for example, when β is monotonically increased (β′≧0), it is possible to prevent the generation of false images in areas where Sorg or Sus is small (low brightness area). However, it has been difficult to prevent the generation of black linear false images on the muscle side at the boundary between bone and muscle. That is, in the conventional method,
False images in which the low-brightness side of edge boundaries in bone and muscle photography is emphasized and becomes less than the fog density of the final recording medium and appears white, or conversely, the high-brightness side becomes too dense and becomes a black line. In addition, it is difficult to completely prevent the occurrence of false images such as false images in which the barium-filled area has a double contour shape in gastric double contrast imaging, and it is difficult to sufficiently improve diagnostic performance. In some cases, this could even lead to misdiagnosis. In order to solve the above-mentioned problems, the present invention has the following points:
It is an object of the present invention to provide a radiation image processing method and apparatus that prevent the generation of false images. That is, an object of the present invention is to provide a radiation image processing method and apparatus that can economically and rapidly obtain radiation images with improved diagnostic performance and without false images. In order to achieve the above object, the present inventor has conducted intensive research and found that the above false image is a difference signal |Sorg-
Sus | is likely to occur in a large area, and based on this knowledge, when processing image signals, S' = Sorg + F (X) (1) (where X = Sorg - Sus, F (X ) is |X ₁ |<
When |X ₂ |, F′(X ₁ )≧F′(X ₂ )≧0, and at least a certain value of X X ₀ (|X ₁ |<|X ₀ |<|X ₂
|) is a monotonically increasing function with F′(X ₁ ) > F′(X ₂ ) as the boundary. The above objective is achieved by reducing the degree of increase in . The radiation image processing method and apparatus of the present invention scans a phosphor with excitation light, reads radiation image information recorded therein, converts it into an electric signal, and then reproduces it on a recording material. , the unsharp mask signal Sus corresponding to the very low frequency at each scanning point
When the original image signal read from the phosphor is Sorg and the reproduced image signal is S', the above calculation formula (1) S'=Sorg+F(X) (X=Sorg-Sus, F(X ) is |X ₁ |<
When |X ₂ |, F′(X ₁ )≧F′(X ₂ )≧0, and at least a certain value of X X ₀ (|X ₁ |<|X ₀ |<|X _{2
| _} It is something to do. Here, |X ₀ | is the difference signal |X|=|Sorg
Needless to say, it is set within the range of -Sus|. In other words, the above F(X) is when X<0
It includes a function in which F″(X)>0, when X>0, F″(X)<0, and a function in which part or all of this curved function is approximated by one or more linear functions. . Moreover, this F(X) is at least X=Sorg
−Sus functions are sufficient, and at the same time Sorg and /
Alternatively, it may be a function of Sus. Mathematically, this means that F(X) in the above equation (1) is β
This only means that cases where it can be replaced with (Sorg).f(x) or β(Sus).f(x) are included. In such a case, the above F′(X),
F″(X) are respectively ∂F(X)/∂X, ∂ ² F
Needless to say, it refers to (X)/∂X ² . The monotonically increasing function F(X)(X
=Sorg−Sus), for example, F(X)=α・sgn(X)・|X| ⁿ +b (2) (where α, b are constants, α>0, 0<n<1,
sgnX=1(X>0), sgnX=-1(X<0),
sgnX=0 (X=0)) F(X)=α・sin(pX) (3) (However, α>0, |pX|<π/2) or F(X)=1−e ^-x (X >0) A curved monotonically increasing function such as F(X)=-1+e ^x (X<0) (4) can be considered. These equations (2), (3), and (4) all have |X ₁ |<|
When X ₂ |, F'(X ₁ )>F'(X ₂ )>0, and the condition F'(X ₁ )≧F'(X ₂ )≧0 is satisfied. Also, all of these functions are F″(X)<0(X>
0), F'' (X) > 0 (X < 0), and in the region where X is positive, as X increases
F'(X) gradually becomes smaller (gradient decreases) and becomes a curved function. However, as mentioned above, as F(X) of the present invention, the second derivative F''(X) of this X may be zero in a part, that is, there may be a straight line part, or it may be zero throughout. It is sufficient that the above conditions are satisfied even if F″(X)
is zero and F′(X ₁ )>F′(X ₂ ), for example, each of the curve-type functions (2) as a whole,
(3) and (4) may be approximated by a broken line that is a combination of multiple linear functions. An example of such a case is, for example (However, a>b>d>0, c=a|X ₁ |-b
｜X ₁ ｜, e=b｜X ₂ ｜+c−d｜X ₂ ｜=a｜X ₁
A polylinear function such as |+b(|X ₂ |−|X ₁ |)−d|X ₂ |) can be considered. Note that all of the F(X)s exemplified above are point symmetrical at the origin, but the F(X)s in the present invention are not necessarily limited to such symmetrical ones at the origin. According to the present invention, the first derivative of F'(X), that is, the function of the points Sorg and Sus, is made smaller as |X| becomes larger, so that the rate of increase in frequency emphasis is suppressed when |X| is large. This makes it possible to increase the degree of frequency emphasis as the difference signal is larger, while suppressing an increase in the degree of emphasis where the difference signal is large, thereby preventing the generation of false images. That is, according to the present invention, frequency enhancement by non-sharp mask processing is normally performed in areas where the difference signal is small in a radiographic image, and in areas where the difference signal is large (for example, the boundary between bone and muscle, the boundary between soft tissue and gas area, and the stomach). The increase in the degree of frequency enhancement is suppressed and the generation of false images is prevented in the border angiography between the Ba-filled area and its surroundings (such as blood vessel shadows in angiography). In particular, F(x) in the above equation (1) is β
If it can be replaced with (Sorg)・f(x) or β(Sus)・f(x), β(Sorg) or β
(Sus) naturally changes depending on Sorg or Sus, so in addition to the above effect,
Needless to say, this also has the effect of making β variable as disclosed in No. 163571. In the present invention, the unsharp mask signal Sus corresponding to ultra-low frequencies refers to an unsharp image (hereinafter referred to as an "unsharp mask") obtained by blurring the original image so that it only contains frequency components lower than the ultra-low frequency components.
refers to the signal at each scanning point. This non-sharp mask has a modulation transfer function of 0.5 or more at a spatial frequency of 0.01 cycles/mm, and 0.5 cycles/mm.
A value that is 0.5 or less at a spatial frequency of mm is used. Furthermore, as a non-sharp mask,
It is preferable to use a non-sharp mask whose modulation transfer function is greater than or equal to 0.5 at a spatial frequency of 0.02 cycles/mm and less than or equal to 0.5 at a spatial frequency of 0.15 cycles/mm, as this significantly improves diagnostic performance. . If the spatial frequency at which the modulation transfer function is 0.5 is c, then in the non-sharp mask used in the present invention, c is 0.01 to 0.5 cycles/mm, preferably
It can be said that it is within the range of 0.02 to 0.15 cycles/mm. In the present invention, the original signal refers to a signal that has been processed by means commonly used in the optical industry, that is, a signal that has been subjected to nonlinear amplification such as logarithmic amplification for band compression and nonlinear correction. Needless to say, it includes. In addition, the method for creating a non-sharp mask is as follows: (1) The original image signal at each scanning point is stored, and depending on the size of the non-sharp mask, the average value (simple average or various (A weighted average of the weighted average of (2) After reading out the original image signal using a small diameter light beam, etc. (3) A method for reading out signals at each scanning point by averaging them together with the surrounding signals using a large diameter light beam that matches the size of the unsharp mask when accumulated images still remain. This method takes advantage of the fact that the beam diameter gradually expands due to scattering of the light beam in the phosphor layer.The original image signal Sorg is created from the light emission signal from the light beam incident side, and the light emission from the side through which the light beam passes with unsharp mask signal
How to create Sus (in this case, the size of the non-sharp mask can be controlled by changing the degree of light scattering of the phosphor layer and the size of the aperture that receives this light), etc. can be used. Among these unsharp mask creation methods,
From the viewpoint of providing flexibility in image processing, method (1) is most preferable. In order to carry out method (1), the following calculations are ideally required to obtain the unsharp mask signal Sus at each scanning point. Here, i and j are the coordinates of a circular area centered on each scanning point (the number of pixels falling within that area is N in the diametrical direction), and aij is a weighting coefficient that is equal in all directions. It is preferable to have directional and smooth changes.

[Formula]. However, if such an operation were to be performed simply, it would be necessary to perform approximately π/4N ² multiplications and π/4N ² additions for each scan point, and if N is large, the operation becomes extremely difficult. The disadvantage is that it is time consuming and impractical. In fact, when reading out a normal radiographic image by scanning a phosphor, it is necessary to avoid losing the frequency components of the image. Although there are some differences, it is usually necessary to scan at a sampling rate of about 5 to 20 pixels/mm (200 to 50μ in terms of pixel size), whereas the non-sharp mask in the present invention is compatible with ultra-low frequencies. Therefore, in order to create this mask, it is necessary to perform calculations using an extremely large number of pixels. For example, for a mask with Gaussian weighting coefficients, if the pixel size is 100μ x 100μ, then if c = 0.1 cycles/mm, N will be approximately 50, and if c = 0.02 cycles/mm, N will be approximately 50. Since the number is approximately 250, the calculation time becomes enormous. Furthermore, averaging a circular area means changing the addition range for each scanning line, but having to make such a judgment during calculation greatly complicates the calculation mechanism. It is uneconomical. In order to solve this problem and perform image processing practically, the method of obtaining a non-sharp mask signal is to use two sides parallel to the main scanning direction and two sides parallel to the sub-scanning direction. The original image signal Sorg at each scanning point within the rectangular area surrounded by
By simply averaging the unsharp mask signal for very low spatial frequencies at each scanning point,
How to obtain Sus (Patent application filed in 1983 by the same applicant)
151400), or in the main scanning direction, the analog signal is de-sharpened using a low-pass filter with a certain reduction transmission characteristic, and in the sub-scanning direction, an A/D
It is preferable to use a method (Japanese Patent Application No. 151402/1989 filed by the same applicant) in which an unsharp mask signal Sus for an ultra-low spatial frequency at each scanning point is obtained by averaging the converted digital signals. In the former case, the weight is uniform in a rectangular area, and therefore the transfer characteristic may oscillate or change in direction, compared to, for example, a mask with a Gaussian distribution weight whose weight decays smoothly. The present inventor has found that there is no substantial difference in terms of improvement in diagnostic performance from the case of the ideal mask calculation described above, despite the drawbacks such as a difference in the degree of unsharpness. It's being served. Moreover, this method uses a rectangular unsharp mask and obtains the unsharp mask signal Sus by simple averaging of the density within the mask.
The unsharp mask signal Sus can be obtained using a very simple method.
can be obtained, resulting in a significant reduction in calculation time and equipment cost. This is an advantage common to both digital and analog signal processing. The transfer characteristic of a rectangular unsharp mask with uniform weights is expressed by the sinc function (sinc(x)=sinπx/πx
), so the above-mentioned provision that the spatial frequency at which the modulation transfer function is 0.5 in the present invention is in the range of 0.01 to 0.5 cycles/mm, preferably 0.02 to 0.15 cycles/mm, does not apply to theory in this case. This is synonymous with setting the length of one side of the rectangular non-sharp mask to 60 mm to 1.2 mm, preferably 30 mm to 4 mm.
Note that even if the shape of the non-sharp mask is rectangular,
It is sufficient that the length of each side is within the above range; for example, a rectangular mask with a large aspect ratio is effective for image processing of linear tomography. In addition, even in the latter method of using a low-pass filter, it is made with a low-pass filter with spatially asymmetric transmission characteristics in the main scanning direction, and digital averaging is performed in the sub-scanning direction, making it irregular in the rectangular area. Although it is based on calculations with a certain weight, there is no substantial difference in terms of improved diagnostic performance from the ideal mask calculation described above, and moreover, the main scanning direction is a low-pass filter. Therefore, it has been found that addition operations using digital signals, which take a long calculation time, can be significantly reduced, resulting in a significant reduction in calculation time and equipment cost. Furthermore, in the latter case, if the average of digital signals in the sub-scanning direction is a simple average, there is no need to perform multiplication, which simplifies the device and speeds up the calculation. The inventor has found that there is virtually no difference in diagnostic performance compared to the ideal case. In the present invention, in addition to the above operations, smoothing processing can also be performed. Generally, in the wavenumber region of ultra-low spatial wavenumbers or higher, there is a lot of noise and it is often difficult to see, so it is often preferable to perform further smoothing processing to further improve diagnostic performance. The smoothing process is preferably such that the modulation transfer function is 0.5 or more when the spatial frequency is 0.5 cycles/mm, and 0.5 or less when the spatial frequency is 5 cycles/mm. What kind of smoothing processing is preferable? For example, when interpreting shadows with a relatively low frequency such as a chest tomogram, it is preferable to remove as much noise as possible. When it is necessary to trace fine blood vessel shadows containing high frequency components, excessively strong smoothing processing may make it difficult to see the desired shadows, which may be undesirable. However, according to the research of the present inventor,
It has been found that performing the smoothing process as described above has the effect of improving diagnostic performance for almost all X-ray images. Furthermore, this smoothing process is equally effective whether it is applied to S′ after the ultra-low spatial frequency processing of the present invention or whether it is applied to the original image signal Sorg. It is recognized that Further, in the present invention, gradation processing may be performed in addition to frequency emphasis processing using a non-sharp mask. Ultra-low frequency processing has a relatively small effect on diseases in which luminance changes slowly over a large area, such as lung cancer and breast cancer. ,same
It is desirable to use the gradation processing disclosed in No. 54-23091, No. 54-23092, etc. in combination. In this case, gradation processing may be performed either before or after ultra-low frequency processing. In the present invention, a phosphor is a material that is irradiated with first light or high-energy radiation, and then stimulated (excited) by optical, thermal, mechanical, chemical, electrical, etc. A phosphor that exhibits so-called photostimulability, which re-emits light corresponding to the irradiation amount of 300~
Those having a stimulable emission wavelength of 500 nm are preferable, such as rare earth element-activated alkaline earth metal fluorohalide phosphors [specifically, patent applications filed in 1973-
84742 (Ba _1-xy ,
Mg _x , Ca _y ) FX: aEu ²⁺ (However, X is at least one of Cl and Br, and x and y are 0<
x+y≦0.6 and xy≠0, and a is 10 ^-6 ≦a≦
(Ba 1 ^- _x , M〓 _x ) FX: yA (however, M
〓 is at least one of Mg, Ca, Cr, Zn and Cd, X is at least one of Cl, Br and I, A is Eu, Tb, Ce, Tm, Dy, Pr,
at least one of Ho, Nd, Yb and Er;
x is 0≦x≦0.6, y is 0≦y≦0.2)
ZnS: Cu, Pb, BaO・xAl ₂ O ₃ : Eu (however, 0.8≦
x≦10) and M〓O・xSiO ₂ :A (however, M〓
is Mg, Ca, Sr, Nn, Cd or Ba, and A is
Ce, Tb, Eu, Tm, Pb, Tl, Bi or Mn, and x is 0.5≦x≦2.5); and patent application No. 1973
−LnOX described in specification No. 84743: xA (however,
Ln is at least one of La, Y, Gd and Lu
1, X is at least one of Cl and Br,
A is at least one of Ce and Tb, x is 0<x<0.1); and the like. Among these, rare earth element-activated alkaline earth metal fluorohalide phosphors are preferred, and among these, barium fluorohalides shown as specific examples are particularly preferred because of their excellent stimulable luminescence. Furthermore, if the phosphor layer of a stimulable phosphor plate made using this stimulable phosphor is colored with a pigment or dye, the sharpness of the final image will be improved and favorable results will be obtained ( Special application 1971-71604
issue). In the present invention, a laser beam with good directivity is used as excitation light for reading out the radiation image accumulated on the stimulable phosphor plate. As an excitation light source for the laser beam, one that emits light in the wavelength range of 500 to 800 nm, preferably 600 to 700 nm, such as He
-Ne laser (633 nm) and Kr laser (647 nm) are preferred, but excitation light sources other than those mentioned above can also be used if a filter that cuts out light other than 500 to 800 nm is used in combination. A radiographic image subjected to image processing according to the present invention is reproduced on a recording medium, and as the recording medium, in addition to silver halide photographic film, diazo film, electrophotographic material, etc. can be used. Further, it may be displayed on a CRT or the like for observation, or it may be optically recorded on a recording material. Hereinafter, the present invention will be explained in detail based on an X-ray photography system as an embodiment thereof. FIG. 1 shows an example of an apparatus for performing image processing according to the present invention in the process of drawing an X-ray photograph. When X-rays are emitted and irradiated onto a human body, the X-rays that pass through the human body enter the phosphor plate. This phosphor plate stores the energy of the x-ray image at the phosphor trap level. A stimulable phosphor sheet 1 on which a radiation image has been accumulated and recorded by this X-ray photography is sent by a roller 2. This roller 2 feeds the sheet 1 in the direction of arrow A and performs sub-scanning for image reading from the sheet 1. Main scanning is performed by scanning a laser beam from a laser light source 3 of excitation light having a wavelength of 500 to 800 nm in the direction of arrow B with a scanning mirror 3a. 300 to 500 nm by scanning this excitation light.
Stimulated light emission in the wavelength range occurs, and this stimulated light is collected by a light condenser 4a made of a light-guiding sheet material, and the light is emitted from a photomal or the like arranged at the output end of this light condenser 4a. It is detected by the detector 4 and converted into an electrical signal. This electrical signal is amplified by an amplifier 5, converted to a digital signal by an A/D converter 6, and sent to a calculation section 7. In the calculation section 7, a calculation device 8a for obtaining the unsharp mask signal Sus
Sus is determined, and then in the difference signal calculation device 8b.
Sorg−Sus is obtained, and then F
After calculating (X), use the above-mentioned calculation formula (1), S′=
This calculation is performed in the calculation device 8d that calculates Sorg+F(X), and the digital signal S′ obtained after the calculation is converted into an analog signal by the D/A converter 9, and after being amplified by the amplifier 10, The light is input to the recording light source 11. Light generated from this recording light source 11 passes through a lens 12 and is irradiated onto a recording material 13, such as a photographic film, mounted on a printing drum 14. A radiographic image is reproduced on this photographic film, and a diagnosis is made by observing this image. The image processing described above may be performed online by directly using the output of the photodetector 4 as in the embodiment described above, or may be performed offline based on data once recorded on a magnetic tape or the like. Unsharp mask processing uses the unsharp mask signal Sus
Using the original image signal Sorg obtained by the photodetector, the calculation is performed as follows: S'=Sorg+F(X) (where F(X) is defined by the above equation (1)). This unsharp mask signal Sus is obtained by the method described later, and it is determined that the modulation transfer function is 0.5 or more at a spatial frequency of 0.01 cycle/mm, and 0.5
Use a modulation transfer function that is less than or equal to 0.5 at a spatial frequency of cycles/mm, or preferably 0.5 at a spatial frequency of 0.02 cycles/mm.
You must specify something that is above and below 0.5 at a spatial frequency of 0.15 cycles/mm. Furthermore, when calculating the above equation, the function F(X) must be specified. This function may be specified individually from the outside, or several types may be determined depending on the part of the human body or each case, and these functions may be stored in the memory of the computing device. Hereinafter, an embodiment in which image processing is performed by specifically determining the F(X) will be described in detail. As mentioned above, F(X) can be changed in various ways, and an appropriate one can be selected from among them, but even if F(X) is not expressed in advance as a general formula, From the value of F
(X) may be obtained to determine F(X). That is, for example, a conversion table between X and F(X) is created on a disk or memory, and the value of F(X) is output according to the input value of X using this conversion table. Good too. In the description of the embodiments below, F(X) is calculated using the function of one aspect of formula (2) as F(X)=α√|| (6), and → A method using a conversion table of F(X) will be explained as a typical example. Figure 2 shows a flowchart for calculating F(X) = α√| An unsharp mask signal (Sus) is calculated and obtained by one of the above-mentioned methods based on 22. Next, using this Sorg and Sus, perform the calculation of X=Sorg-Sus to find X23. When X is positive or 0, replace X with X and α with α, 24, 25, and when X is negative, replace
Sorg+F(X), ie, Sorg+α√, is calculated to obtain 27, S′. To perform the calculation shown in the flowchart of FIG. 2, an input/output device 31, a controller 32, a calculation unit 33, and a memory 34 are used connected to the data bus 30 as shown in FIG. Vessel 33
must have the ability to calculate square roots (√) and perform addition, subtraction, multiplication, and division. Figure 4 shows X→F(X) (for example, F(X)=α√
) shows an example of using the conversion table of
In this case, Sorg and Sus are found 41, 42 in the same way as the example shown in FIG. 2, and X is found by calculating X=Sorg-Sus 43. From the X obtained here, convert the data by referring to the conversion table (X → F (X)),
Find F(X) according to X. 44. Next, using F(X) obtained in this way, Sorg+F
Perform the calculation of (X) 45 and find S'. In order to perform the calculation shown in the flowchart of FIG. 4, in addition to using an input/output device 51, a controller 52, a calculation unit 53, and a memory 55 connected to a data bus 50 as shown in FIG. Memory 54 is used. In this case, the computing unit 53 is F
Since there is no need to calculate (X)=α√||, any method that can perform addition and subtraction is sufficient. Furthermore, by performing smoothing processing for reducing high frequency components on the frequency-emphasized signal S' as described above, noise can be reduced without damaging information necessary for diagnosis. Furthermore, in addition to frequency emphasis processing using a non-sharp mask, gradation processing can also be used in combination. When gradation processing is performed before ultra-low frequency processing, A/D conversion is performed after gradation processing is performed using a nonlinear analog circuit. If it is performed after A/D conversion, digital processing can also be performed using a minicomputer. Further, digital processing is performed after ultra-low frequency processing, or analog processing is performed after D/A conversion. Also, when playing back and recording images on photo film,
A reduced photographic image can be obtained by recording at a higher sampling frequency than during input scanning. For example, in the input system, 10
Pixels/mm, if you scan at 20 pixels/mm in the output system, it will be 1/
The photographic image is reduced to 2. Like this 1/2~1/
In a photographic image reduced to 3.3, the frequency components considered necessary for diagnosis are close to the frequency range with the highest visual sensitivity, so the contrast appears to be higher visually, making it much easier to see. Note that in the above embodiment, the original image signal Sorg may also mean a signal that has been subjected to band compression such as logarithmic transformation and nonlinear correction.
Practically speaking, since the output of the photodetector is subjected to signal processing, it is desirable to perform band compression such as logarithmic transformation. In principle, the output of the photodetector can be used as is.
Needless to say, it is also possible to perform subsequent processing as Sorg. In addition, theoretically, when calculating this mask, the average energy should be calculated, but according to the inventor's experiments, when calculating this unsharp mask signal, a value corresponding to the logarithmically compressed density is used. Even if I took the average value, the results remained the same. This is practically advantageous in terms of processing. Examples will be given below to make the results of the present invention more clear. Examples A total of 50 cases of the typical sites listed in Table 1 were recorded directly on conventional X-ray film and read out from the phosphor according to the present invention, and are shown in Fig. 6 as F(X). Curves A, B, D
and polygonal line C, and compare them with a photographic image created by applying ultra-low frequency processing according to the above calculation formula (1),
We investigated the improvement of diagnostic performance for major parts of the human body. Solid line A in Figure 6 is F(X) = 0.4・sgn(X) |
X|〓, that is, α= in the above function (2)
0.4, n=1/2, and b=0, and is a continuous curve-type function. The gradient (F′(X ₁ )) is |X|
has become smaller as the number of
F″(X) is negative when X is positive, and positive when X is negative. Broken line B is F(X)=1−e ^-1 . ^4x (X>0), F
(X)=-1+ ^e1.4x (X<0) ^, i.e. the above function
In (4), the coefficient of X is set to 1.4, which is also a continuous curve-type function. Gradient (F′(X))
becomes smaller as |X| increases,
F″(X) is negative when X is positive and positive when X is negative. The chain line C is F(X)=sgn(X)(n|X|+
This is a polygonal function in which n of const) is made smaller depending on the size of |X|, and the larger |X|
In the above function (5), a=1, b=0.75, c=
Corresponds to 0.5. That is, F(X)=X |X ₁ | <0.1 F(X)=sgn(X) (0.75 | |X|+0.1) 0.3≦|X| In this function, the gradient F′(X) gradually decreases as |X| increases, and F″(X) is 0. The dotted line D becomes convex upward in the region where X>0. Nari, X
A conversion table that is convex downward in the area <0 is created. This is an example in which the shape is not symmetrical in the region of X>0 and X<0. In this case, calculations were performed using the Table Lutzking method. Regarding the existence and degree of improvement in diagnostic performance, the physical evaluation values of ordinary photography (for example,
Since it is virtually impossible to substantiate the results using information such as sharpness, contrast, graininess, etc., the evaluation was based on subjective evaluations by four radiologists. The evaluation results are shown in Table 1. Table 1 Cases, site evaluation Head: The skull was white and did not come off, and there were no black line-like false images of the facial muscles, making it easy to see and diagnosing muscle tumors. Bone and muscle: Accurate diagnosis was made in both bone and muscle areas, with no false images occurring. Angiography: No false images occur, and it is now possible to diagnose both thin and large contrast-enhanced blood vessels. Double contrast imaging of the stomach: No false images occur in the peripheral region of the stomach and areas filled with a large amount of contrast medium.
The overall diagnosis was good. Simple abdomen: The gas area of the intestines is not emphasized more than necessary, making it easier to diagnose the entire abdomen. In addition, as F(X), curves A, B, and
Although some differences in the degree of improvement in diagnostic performance were observed for individual photographic images depending on which of the lines D and C was selected, on average no substantial difference was observed for each case. As is clear from Table 1, according to the present invention, generation of false images was prevented in various cases and regions, and diagnostic performance was improved. As described above, the method of the present invention can greatly improve diagnostic performance in a radiation image processing method using a stimulable phosphor, and has a remarkable practical effect.

[Brief explanation of the drawing]

FIG. 1 is a schematic block diagram showing an example of a radiation image recording and reproducing system implementing the radiation image processing method of the present invention, FIG. 2 is a flowchart showing an example of the method of the invention, and FIG. 3 is similar to FIG. 4 is a flowchart showing another example of the method of the present invention, and FIG. 5 is a block diagram showing an example of the configuration of a calculation unit used to implement the method shown in FIG. Fig. 6 is a block diagram showing an example of the configuration of the arithmetic unit used in the present invention, and Fig. 6 is a graph showing each example of the function F(X) of the arithmetic expression S'=Sorg+F(X) used in the method and apparatus of the present invention. be. DESCRIPTION OF SYMBOLS 1...Stormable phosphor sheet, 3...Laser light source, 4...Photodetector, 4a...Concentrator, 5, 10
...Amplifier, 7...Arithmetic section.

Claims (1)

[Claims] 1. Scanning a stimulable phosphor material to read radiation image information recorded in the phosphor material and converting it into an electrical signal, and then reproducing it as a visible image on the recording material. Then, the unsharp mask signal Sus corresponding to the ultra-low spatial frequency at each scanning point is determined, and the original image signal read out from the phosphor is converted to Sorg,
When the reproduced image signal is S', the calculation formula S'=Sorg+F(X) (However, X=Sorg-Sus, F(X) is |X ₁ |<
When |X ₂ |, F′(X ₁ )≧F′(X ₂ )≧0, and at least
Performs _{the operation expressed by a} certain _value A radiation image processing device characterized by emphasizing frequency components higher than the ultra-low spatial frequency. 2. A patent characterized by using a non-sharp mask whose modulation transfer function is 0.5 or more at a spatial frequency of 0.01 cycles/mm and 0.5 or less at a spatial frequency of 0.5 cycles/mm. A radiation image processing method according to claim 1. 3. A patent characterized in that a non-sharp mask is used which has a modulation transfer function of 0.5 or more at a spatial frequency of 0.02 cycles/mm and 0.5 or less at a spatial frequency of 0.15 cycles/mm. A radiation image processing method according to claim 1. 4. A patent characterized in that the monotonically increasing function F(X) is a curved function that satisfies the following conditions: F″(X)<0 (X>0) F″(X)>0 (X<0) A radiation image processing method according to any one of claims 1 to 3. 5 The monotonically increasing function F(X) mainly consists of a curved part that satisfies the following conditions: F″(X)<0 (X>0) F″(X)>0 (X<0) Claims 1 to 3 include a linear portion where ``(X)=0.
3. The radiation image processing method according to any one of paragraphs. 6 A plurality of linear functions that approximate a curved function where the monotonically increasing function F(X) satisfies the following conditions: F″(X)<0 (X>0) F″(X)>0 (X<0) A radiation image processing method according to any one of claims 1 to 3, characterized in that the method comprises a combination of the following. 7 The function F(X) is F(X)=α・sgn(X)・|X| ⁿ +b (However, α and b are constants, α>0, 0<n<1 sgn(X)=1 X> 5. The radiation image processing method according to claim 4, wherein the radiation image processing method is a curve-type function expressed as follows. 8. Claims characterized in that the function F(X) is a curved function expressed as F(X)=α・sin(pX) (where |pX|<π/2, α>0) 4. The radiation image processing method according to item 4. 9. A patent characterized in that the function F(X) is a curved function expressed by F(X)=1-e ^-x (X>0) F(X)=-1+e ^x (X<0) A radiation image processing method according to claim 4. 10 An excitation light source for scanning a stimulable phosphor to stimulate the radiation image stored and recorded therein, a photodetector for detecting this emission and converting it into an electrical signal, and a photodetector for detecting this emission and converting it into an electrical signal. In a signal processing device in a radiographic image recording and reproducing system, which is equipped with a processing arithmetic unit, the arithmetic unit processes the detected original image signal as Sorg and the unsharp mask signal corresponding to the ultra-low spatial frequency at each detection point as Sus. When, S′=Sorg+F(X) (However, X=Sorg−Sus, F(X) is |X ₁ |<
When |X ₂ |, F′(X ₁ )≧F′(X ₂ )≧0, and at least a certain value of X X ₀ (|X ₁ |<|X ₀ |<|X ₂
A radiation image processing device characterized in that it performs an operation expressed by a monotonically increasing function (F′(X ₁ )>F′(X ₂ ) with |) as the boundary.